Many applications involve taking more than one type of measurement of electromagnetic radiation. For example, some medical imaging applications involve insertion of an endoscope into a cavity or incision in a subject such as a human patient. A flexible endoscope, for example, may include an optics channel through which a first optical fiber bundle conveys illumination light for illuminating internal tissues of the patient, and through which a second, coherent optical fiber bundle conveys light reflected or fluorescently emitted by the internal tissues back up through the endoscope to a measuring device such as a charge-coupled device (CCD) camera. A resulting image of the internal tissues produced by the camera may then be displayed on a monitor for visual inspection by a surgeon or physician, who may be able to identify suspected abnormal or diseased tissue from the displayed image.
Once suspected abnormal tissues have been identified by such visual inspection, it is then desirable to perform additional analysis on the tissue to confirm with greater specificity or accuracy whether it is in fact diseased. For this purpose, spectroscopy is sometimes performed. One existing spectroscopic analysis method involves the insertion of an optical fiber probe through a biopsy channel of the endoscope, which is normally used for insertion through the endoscope of medical tools such as those used for tissue sampling or therapeutic interventions, for example. The presence of this optical fiber probe in the biopsy channel may make it difficult or impossible to insert other tools into the biopsy channel, rendering the biopsy channel unsuitable for its intended purpose. In addition, this procedure may pose inconvenience for the surgeon or physician, who may have to remove medical tools from the biopsy channel in order to insert the optical fiber probe, then remove the probe in order to reinsert the tools when the spectral measurement is completed. Moreover, when the optical fiber probe is inserted through the biopsy channel, the probe typically comes into physical contact with the tissue in order to perform a measurement. Such contact tends to press blood away from the tissue to varying degrees, depending on the amount of pressure applied, which may result in different observed spectra, thereby introducing a source of measurement error.
One existing endoscopic system employs a beam splitter for directing a percentage of radiation received from the tissue for receipt by a spectroscopy device, while allowing the remainder of such radiation to pass through the beam splitter for receipt by a camera. However, it will be appreciated that beam splitters of this nature reduce the intensity of light received across the entire area of the camera. Generally, only a relatively low amount of light from the analyzed tissues enters the endoscope, due to the small circumference of the endoscope, the limited ability to increase the intensity of the illuminating light without causing thermal damage or photobleaching in the tissue, and due to the relatively low intensity of light fluorescently emitted or reflected by the tissue. Accordingly, the CCD camera is already “light hungry”. The use of such beam splitters aggravates this problem, resulting in an even darker CCD image, which may necessitate the use of expensive signal amplification devices.
Alternatively, in another existing endoscopic system, a mirror is employed for a somewhat different purpose. The mirror is inserted into the optical path of the light beam from the endoscope so as to reflect the entire beam to a first camera for white light reflectance imaging, and is removed from the optical path so as to allow the entire beam to be received at a second camera for fluorescence imaging. However, this method does not allow for simultaneous measurements by the first and second cameras, which increases the chance that the endoscope or the subject might move between alternate images. This difficulty may not be serious for use in switching between white light reflectance and fluorescence images, however, this method would not be desirable for combined imaging and spectroscopy measurements, as it fails to ensure that the spectroscopy measurement is of the same tissue region that appeared to be of interest in the camera image, which may lead to unreliable spectroscopy results.
Accordingly, there is a need for a more convenient way of performing contemporaneous measurements with multiple measuring devices, such as an imaging device and a spectroscopy device for example, without significantly compromising endoscopic imaging quality or reliability of the spectroscopy results.
Additionally, existing endoscopic systems have failed to utilize the full potential of combined imaging and spectroscopy. In particular, for systems involving multi-spectral-channel imaging devices, such as white light reflectance RGB color CCD cameras and dual channel fluorescence imaging cameras for example, the ability to increase the diagnostic sensitivity of such devices by adjusting the gain relationships between different imaging channels is constrained by conventional wisdom, which indicates that any increase in the diagnostic sensitivity of the imaging device by gain relationship adjustment results in a corresponding decrease in specificity of diagnosis. In other words, increasing the diagnostic sensitivity of a dual channel fluorescence imaging device, for example, will produce more “false positive” diagnoses, as a result of tissues that appear from the image alone to be diseased or malignant when in fact they are benign or even normal. The desire to avoid such erroneous diagnoses therefore places limitations on the ability to adjust the diagnostic sensitivity of the imaging device.
Thus, there is a need for a way to produce images of higher diagnostic sensitivity without unduly reducing the specificity or accuracy of diagnoses.
In addition, an endoscopic imaging system preferably involves both white light reflectance color imaging to produce a normal view in which the appearance of an internal organ is familiar to the surgeon or physician, and fluorescence imaging for better diagnostic accuracy. For white light reflectance imaging, an image of the tissue of interest is taken while the tissue is being irradiated with white light. For fluorescence imaging, the tissue is irradiated with excitation light, typically short wavelength light, which may range from blue to ultraviolet depending on the application. In order to avoid the necessity of injecting the tissue with drugs containing fluorescent substances, the trend has been toward autofluorescence imaging. When tissues are irradiated with short wavelength excitation radiation, the tissues tend to emit fluorescence light which typically ranges from 450 to 750 nm and peaks at green wavelengths from 510 to 530 nm, for example. It has been found that abnormal tissues such as diseased or cancerous tissues tend to emit significantly lower intensities of such autofluorescence light at green wavelengths than normal tissues. Abnormal or suspicious tissues therefore tend to appear darker in a corresponding fluorescence image of the tissues at green wavelengths. Thus, different illumination spectra are required for reflectance and fluorescence imaging, namely, a white light or other illumination spectrum for reflectance imaging and at least a short-wavelength excitation spectrum for fluorescence imaging.
Most existing systems for reflectance and fluorescence imaging are either inconvenient to switch between reflectance and fluorescence imaging, or fail to adequately correct the fluorescence image to compensate for geometric factors, or both.
More particularly, to switch between white light reflectance and fluorescence imaging, many systems require a user of the system, such as a surgeon or physician, to manually disconnect a first light source and first RGB CCD camera used for white light reflectance imaging from the endoscope, and to connect a second separate light source and second fluorescence camera to the endoscope for fluorescence imaging. Such manual disconnection and connection of light sources and cameras are time-consuming and inconvenient to the user, and increase the duration and discomfort to the patient being examined.
With respect to correction of the fluorescence image to compensate for geometric factors, it has been found that using only a single short-wavelength illumination waveband is disadvantageous for fluorescence imaging. Although tissue abnormality or disease may cause a given point in the fluorescence image to appear dark, alternatively, normal tissue may also appear dark if it is simply further away from the tip of the endoscope than other points in the tissue, or alternatively normal tissue may appear dark due to partial obstruction or other geometrical factors, such as curved tissue surfaces, folds, polyps, or the angle of the endoscope relative to the tissue surface, for example. Thus, it is not possible to determine from a green fluorescence image alone whether or not a particular point in the tissue appears dark because it is abnormal, or whether it appears dark merely because of its distance or geometrical positioning relative to the endoscope tip.
Some systems have attempted to address the latter difficulty by additionally measuring autofluorescence at red wavelengths, as autofluorescence intensities of normal and abnormal tissues are more similar at red and longer wavelengths than they are at green wavelengths. The resulting red autofluorescence image may be used to correct the green autofluorescence image for the geometry of the tissue. For example, if the red autofluorescence image is displayed as a red image on a display screen, and the green autofluorescence image is superposed over the red image, then if a given point in the tissue is normal tissue but appears dark in the green image due to geometric factors, then that point will also appear dark in the red image, and will therefore appear dark in the superposition of the two images. However, if a given point in the tissue appears dark in the green image because it is abnormal or diseased, then that point will probably appear bright in the red image, and will therefore appear as a red spot in the superposed image. Unfortunately, however, red autofluorescence occurs at much lower intensities than green autofluorescence, and accordingly, the red image suffers from a low signal-to-noise ratio. In addition, although red autofluorescence emission intensities are similar for normal and abnormal tissues, there is still some difference between the two. Thus, this method tends to suffer from significant measurement error.
One existing system, recently designed in part by some of the inventors of the present invention, has partly addressed both of the above difficulties. An arc lamp directs input radiation onto a cold mirror, which reflects near ultraviolet and visible light to an optical system, while transmitting over 90% of infrared (IR) radiation away from the optical system to prevent heat damage of the optical system due to continuous IR irradiation. The radiation from the cold mirror passes through a long wave pass (LP) filter which transmits visible light through the optical system while attenuating ultraviolet wavelengths. The visible light from the LP filter is then directed through one of a plurality of different filters on a rotary filter wheel. One of the filters generates uniform white light for normal reflectance imaging of the tissue. Another of the filters is a notch-band filter for fluorescence imaging. This way, one light source provides illumination for both white light reflectance imaging and fluorescence imaging, eliminating the need to switch the endoscope between two light sources.
The notch-band filter transmits blue wavelengths shorter than 450 nm, and also transmits red wavelengths longer than 590 nm, which also include some IR wavelengths due to the imperfection of the cold mirror. The notch-band filter attenuates green wavelengths between 450 nm and 590 nm, in order to prevent reflection by the tissue of such wavelengths which would interfere with the ability to measure autofluorescence emission by the tissue at these wavelengths. The blue wavelengths excite the tissue resulting in autofluorescence emission by the tissue at the green wavelengths, which may then be measured to produce a green autofluorescence image. The red wavelengths are used to illuminate the tissue to produce a separate red reflectance image of the tissue, simultaneously with the production of the green autofluorescence image. The red reflectance image has much greater intensity than a red autofluorescence image, and therefore has an improved signal-to-noise ratio, thus reducing errors. The red and green images are then superposed on a display, to provide an improved correction for geometric factors.
However, the single optical system light source employed in the above method tends to be inflexible in at least some respects. For example, because both the blue light used for excitation and the red light used for correction must pass through a single notch-band filter, the selection of wavelengths to be used for excitation and correction is limited by manufacturing constraints on such filters. For example, it may be desirable to use NIR radiation rather than red radiation to provide the reflectance image for correction purposes, as diseased and normal tissues exhibit even more similar reflectance intensities at some NIR wavelength bands than at red wavelengths. However, it may not be feasible to design a single filter with a wider notch-band, to attenuate wavelengths from 450 to 750 nm, for example. Simply eliminating the cold mirror and transmitting all infrared wavelengths through the optical system would be undesirable, as it may cause heat damage to other filters on the rotary filter wheel such as the reflectance imaging filter for example, and may also cause such damage to lenses and other components in the optical system.
Thus, in addition to the deficiencies in existing endoscopic imaging and spectroscopy systems referred to above, there is also a need for an improved illumination source suitable for both reflectance and fluorescence imaging.
Similarly, existing cameras for reflectance and fluorescence imaging are often large and heavy due to the significant number of moving parts they contain in order to switch between reflectance and fluorescence imaging. Such cameras are therefore difficult for a physician or surgeon to wield. Thus, there is also a need for an improved, more light-weight and compact camera capable of performing both reflectance imaging and fluorescence imaging without unduly increasing the size and weight of the camera.
Finally, it is known that cancerous tissues exhibit hypoxia, which is caused by increased oxygen consumption due to rapid growth of cancerous cells. However, other unrelated chromophores tend to overwhelm and obscure the effects of hypoxia at visible imaging wavelengths, with the result that conventional endoscopic imaging systems have typically been unable to detect tissue oxygenation status. Accordingly, there is a need for a way to take advantage of this property of cancerous tissues to imp rove diagnostic accuracy in endoscopic imaging systems.